Amperometric sensor and method for its manufacturing

ABSTRACT

An in vivo amperometric sensor is provided for measuring the concentration of an analyte in a body fluid. The sensor comprises a counter electrode and a working electrode, and the working electrode comprises a sensing layer which is generally water permeable and arranged on a support member adjacent to a contact pad. The sensing layer comprises an immobilized enzyme capable of acting catalytically in the presence of the analyte to cause an electrical signal. The sensing layer has an upper surface facing the body fluid and a lower surface facing away from the body fluid, and the immobilized enzyme is distributed within the sensing layer in such a way that the enzyme concentration in the middle between the upper and lower surfaces is at least as high as on the upper surface of the sensing layer.

CLAIM OF PRIORITY

The present application is continuation of PCT/EP 2007/004606, filed May24, 2007, which claims priority to U.S. Provisional Application No.60/805,151, filed Jun. 19, 2006, each of which are hereby incorporatedby reference in their entireties.

TECHNICAL FIELD OF THE INVENTION

The present application relates to implantable sensors, particularly tosuch sensors for in vivo amperometric determination of analyteconcentration, and more particularly to an amperometric sensorconfigured for implantation into the living body of a human or animal tomeasure the concentration of an analyte in a body fluid, comprising acounter electrode and a working electrode that comprises a generallywater permeable sensing layer comprising an immobilized enzyme capableof acting catalytically in the presence of the analyte to cause anelectrical signal.

BACKGROUND

Sensors of the general type discussed in the context of the presentinvention are generally known; see, e.g., EP 0 247 850. Implantablesensors for the in vivo-measurement of medically important analytes likeglucose or lactate are based on electrochemical enzymatic detection ofthe analyte. The most common approach is the use of an oxidase tooxidize an analyte, e.g. glucose, with subsequent reduction of oxygen tohydrogen peroxide and amperometric detection of the hydrogen peroxide bya working electrode of the sensor. Another approach in the field ofin-vivo sensing bypasses the use of oxygen/peroxide as a mediator coupleby employing synthetic redox mediators for glucose conversion withoutoxygen. In that case, synthetic redox mediators are built into thesensing element. An example for this approach utilizespoly(biimidizyl)osmium complexes as redox mediators in conjunction withenzyme, as described by Feldmann et al. in Diabetes Technology andTherapeutics, 5, 769 (2003).

Despite intensive research and development efforts there are at presentno implantable sensors available which measure medically importantanalytes like glucose reliably over extended periods of time.

It is an object of the present invention to provide a way to improve thereliability and longevity of amperometric sensors for invivo-measurements of an analyte concentration in a body fluid.

SUMMARY

According to the invention this object is achieved by an anperometricsensor configured for implantation into the living body of a human oranimal to measure the concentration of an analyte in a body fluid, saidsensor comprising a counter electrode and a working electrode, saidworking electrode comprising a sensing layer which is generallypermeable for water and arranged on a support member adjacent to acontact pad, said sensing layer comprising an immobilized enzyme capableof acting catalytically in the presence of the analyte to cause anelectrical signal, the sensing layer having an upper surface facing thebody fluid and a lower surface facing away from the body fluid, whereinthe immobilized enzyme is distributed in the sensing layer in such a waythat the enzyme concentration in the middle between the upper and lowersurfaces is at least as high as on the upper surface of the sensinglayer.

In a planar configuration, the contact pad can be placed directly underthe sensing layer (or the sensing layer is arranged on the contact pad)both having consistent surface areas. In another embodiment, the contactpad can be made smaller or larger than the sensing layer. In stillanother embodiment, the contact pad can be partly displaced from thearea covered by the sensing layer, so that only a fraction of thesensing layer is in direct contact with the pad. For other arrangements,the contact pad can be placed on one of the sides of the sensing layer.All these options are summarized by the phrase “the sensing layeradjacent to the contact pad”. It is understood that this holds likewisefor the other electrodes.

It has been found that measurements of implanted amperometric sensorsare often adversely affected by a low oxygen concentration insubcutaneous tissue surrounding the sensor. This problem seems to beespecially pronounced in the case of enzymatic sensors which rely on anoxidase (for example a glucose oxidase) as the immobilized enzyme in thesensing layer, because such sensors cause an electrical measurementsignal by oxidizing the analyte. In principle, the strength of themeasurement signal created by such sensors depends on the amount ofenzyme, analyte and oxygen present. If the oxygen concentration issufficiently high, the response of a given sensor with a defined enzymeloading reflects the concentration of the analyte in the vicinity of thesensor and is, in ideal cases, proportional to it. However, if theoxygen concentration is too low, fewer analyte molecules are oxidizedand consequently a weaker electrical signal is produced compared to asensor operating under oxygen-saturated conditions.

Lowering the enzyme loading of a sensor lowers the critical oxygenconcentration at which saturation is achieved, but also reduces thesignal to noise ratio because a smaller measurement signal is created.Hence, lowering the enzyme loading is not sufficient to solve theproblem.

An amperometric sensor according to the present invention deals with theproblem of low oxygen concentrations in subcutaneous tissue by means ofthe sensing layer which comprises the immobilized enzyme which isdistributed in the sensing layer in such a way that the enzymeconcentration in the middle between the upper and lower surface is atleast as high as on the upper surface of the sensing layer.

Accordingly, only a relatively small fraction of the enzyme moleculescontained in the sensing layer is active on the upper surface of thesensing layer. As a consequence, a relatively low concentration ofoxygen is sufficient to saturate the surface of the sensing layer withoxygen. The structure of the sensing layer allows analyte molecules todiffuse into the sensing layer and interact with enzyme moleculesfurther away from the surface which are surrounded by their ownreservoir of oxygen molecules. Thus, the electrical signal of a sensoraccording to the present invention is created not only within a smallsurface layer but rather within an extended volume which lowers theoxygen density (oxygen concentration) at which saturation of the sensoris achieved. Consequently, saturation of enzyme with oxygen can beachieved at lower oxygen concentrations without lowering the signal tonoise ratio of the measurement signal of the sensor.

Amperometric sensors with a porous sensing layer are known in the art.However, the enzyme is supplied to such sensors only after the porouslayer has already been prepared. Therefore, the enzyme concentration ofsuch sensors is typically highest on the upper surface of the sensinglayer and decreases strongly with increasing distance from the surface.As a consequence, the major part of the electrical signal of such asensor is created on the surface of the sensing layer, i.e. in arelatively small volume so that correspondingly higher oxygenconcentrations are required for precise measurements.

A distribution of enzyme throughout the sensing layer according to thepresent invention, especially a homogenous concentration, can be mosteasily achieved by mixing enzyme into a paste, for example a pastecomprising carbon particles and a binder, and applying this mixture ontoa contact pad to provide the sensing layer of a working electrode. Inone embodiment, a surface-active agent such as a detergent or ahydrophilic polymer is used to aid the dispersion of the enzyme withinthe paste. In this way an equal distribution of enzyme moleculesthroughout the sensing layer can be achieved. The object of theinvention is therefore also achieved by a method for manufacturing anamperometric sensor configured for implantation into the living body ofa human or animal to measure the concentration of an analyte in a bodyfluid, said method comprising the following steps: mixing carbonparticles, enzyme and a polymeric binder to create a paste; applyingthat paste adjacent to a contact pad onto a support member; andhardening that paste into a porous sensing layer.

The invention is to be explained in more detail by the following figuresand examples.

BRIEF DESCRIPTION OF THE DRAWINGS

The following detailed description of the embodiments of the presentinvention can be best understood when read in conjunction with thefollowing drawings, where like structure is indicated with likereference numerals and in which:

FIG. 1 shows a first exemplary embodiment of a sensor according to thepresent invention in a cross section view.

FIG. 2 shows a functional characteristic of the sensor according to FIG.1 from an in vitro. measurement.

FIG. 3 shows measurement data of the sensor according to FIG. 1 measuredin a biomatrix.

FIG. 4 shows the dependence of sensor current on a diffusion barriercovering the sensing layer of sensors F to J, but not sensors A to E.

FIG. 5 shows a second exemplary embodiment of a sensor according to thepresent invention in a cross section view.

In order that the present invention may be more readily understood,reference is made to the following detailed descriptions and examples,which are intended to illustrate the present invention, but not limitthe scope thereof.

DETAILED DESCRIPTION OF EMBODIMENTS OF THE PRESENT INVENTION

The following descriptions of the embodiments are merely exemplary innature and are in no way intended to limit the present invention or itsapplication or uses.

FIG. 1 shows schematically a first embodiment of an amperometric sensor1 configured for implantation into the living body of a human or animalto measure the concentration of an analyte in a body fluid of a human oranimal. For better illustration of some details FIG. 1 is not to scale.

In the illustrated embodiment, the sensor 1 comprises a counterelectrode 2, a working electrode 3 and a reference electrode 4 which arearranged on a support member 5 which is typically made of a plasticmaterial, such as polyimide. Each electrode 2, 3, 4 comprises a contactpad 6, 7, 8 which is provided as an electrically conductive film, e.g. ametallic film which can be a gold film, with a thickness of betweenabout 50 nm and about 150 nm. It is also possible to make the contactpads 6, 7, 8 from other metals, in particular palladium, or as amultilayer film of different metals. For example, a thin film, less than20 nm, of titanium covering the support member 5 can be covered with asecond film of gold with a thickness of between about 50 nm and about130 nm, thus forming the contact pads 6, 7, 8. As an alternative, thecontact pads 6, 7, 8 can be formed as an electrically conductive polymerfilm, e.g. from a conductive polymer paste, for example byscreen-printing or by dispensing which leads to thicker contact pads 6,7, 8. Instead of separate counter and reference electrodes 3, 4 acombined counter/reference electrode may also be used. One example of asuitable counter/reference electrode is a silver/silver-chlorideelectrode. As such counter and/or reference electrodes are commonly usedno further description is necessary.

The working electrode 3 further comprises a sensing layer 9 which isgenerally permeable for water and arranged adjacent to the contact pad 7of the working electrode 3. The sensing layer 9 comprises immobilizedenzyme capable of acting catalytically in the presence of the analyte tocause an electrical signal. In the present example an oxidase,particularly a glucose oxidase is used as enzyme to measure glucose asan analyte in a human body fluid, like interstitial fluid or blood.

The sensing layer 9 was applied as a paste onto the support member 5 tocover the contact pad 7 of the working electrode 2. That paste was madeby mixing carbon particles, enzyme and a polymeric binder. In this waythe immobilized enzyme is distributed equally throughout the sensinglayer 9. A homogeneous enzyme distribution throughout the sensing layer9 is advantageous. Hence, the enzyme concentration should differ by lessthan 20%, especially less than 10%, between the upper surface and thelower surface of the sensing layer 9. As the analyte can diffuse intothe porous sensing layer 9 the electrical measurement signal is creatednot just in the sensing layers 9 upper surface which faces away from thecontact pad 7 but rather in an extended volume. Therefore, rather lowoxygen -concentrations are sufficient to saturate the sensor 1 withoxygen and enable precise measurements.

In one embodiment the sensing layer 9 is flat. In other embodiments, thesensing layer 9 is electrically conductive. In yet other embodiments,the electrical conductivity of the sensing layer 9 is at least about 1Ω⁻¹ cm⁻¹. As a result of such conductivity, every place in the sensinglayer 9 at which the enzymatic reaction of the analyte takes places actsas a tiny electrode, at which the product of the enzymatic reaction canbe quickly reduced or oxidized. In this manner these places are used ascathode or anode, depending on the sign of the electrical potentialapplied. Consequently the sensing layer 9 in the porous structurecomprises a huge number of tiny cathodes or anodes. As a result of thisthere is no need for the product of the enzymatic reaction to advancethrough the bulk of the sensing layer 9 which would result in a loss asignal height. An electrically conductive embodiment of the sensinglayer 9 therefore has the result of an increased signal height.

The sensing layer 9 of the example shown has a thickness of about 30 μm.Generally the sensing layer 9 should have a thickness of at least about5 μm, at least 10 μm in one embodiment, in order to provide asufficiently large volume for the creation of the electrical measurementsignal. A thickness of the sensing layer 9 of more than about 100 μmdoes not provide additional benefits. In other embodiments, the sensinglayer 9 has a thickness of between about 20 μm and about 70 μm. In yetother embodiments, the sensing layer 9 is arranged in a depressionprovided in the support member 5. In this way it is somewhat protectedby lateral walls of the depression in the support member 5 from damagesduring the implantation process. Furthermore, the lateral surfaces ofthe sensing layer 9 can be connected to the support member 5 and therebyensure that analyte molecules can diffuse only through the sensinglayer's 9 upper surface into the sensing layer 9. Of course, the lateralsurfaces can be made generally impervious to water by different means aswell. The sensing layer 9 in other embodiments can have lateral surfaceswhich are generally impervious for the body fluid.

In similar fashion, in one embodiment the contact pads 6, 8 of thecounter electrode 2 and the reference electrode 4 are covered withwater-permeable layers 12, 14 which were also applied as paste. Ofcourse, the layers 12, 14 of the counter electrode 2 and of thereference electrode 4 contain no enzyme. Like the sensing layer 9,layers 12 and 14 may also comprise carbon particles and a polymericbinder. Whereas porosity enhancing particles 13 like carbon nanotubeshave been added to the pastes for the sensing layer 9 and the layer 12,such porosity enhancing particles 13 provide no benefit for the highlyconductive layer 14 of the reference electrode 4 and were therefore notadded as a matter of design choice. However, such carbon nanotubes maybe useful for other purposes, as will be described below, and thus theinclusion of them will not detract from the scope of the presentinvention.

As enzyme is distributed throughout the whole sensing layer 9, oxygensaturation can be maintained even if much higher analyte concentrationsare present at the upper surface of the sensing layer 9 than is feasiblefor known sensors. The sensing layer of sensors according to prior artis usually covered by a diffusion barrier which hinders analytediffusion to such an extent that the analyte concentration at the uppersurface is typically about 100 times lower than in body fluidsurrounding the sensor.

The sensing layer 9 of the sensor 1 of one embodiment of the presentinvention is covered by a diffusion barrier which hinders diffusion ofanalyte molecules only to such an extent that, after implantation intothe living body of a human or animal, the analyte concentration at theupper surface of the sensing layer 9 is at most about ten times lowerthan in the body fluid surrounding the implanted sensor 1. In otherembodiments it can be at most about five times lower, and in yet otherembodiments it can be at most about three times lower. In the exampleshown, the diffusion barrier comprises several distinct layers 10, 11contributing to the diffusion resistance of the diffusion barrieragainst diffusion of analyte molecules.

The diffusion barrier is permeable for the analyte and prevents enzymefrom leaking out of the sensing layer 9. In the example shown, thediffusion barrier comprises as a first layer an electrically conductiveenzyme-free layer 10 which comprises carbon particles and a polymericbinder and has a thickness of less than a third of the thickness of thesensing layer 9. Usually it is about 1 μm to about 3 μm thick. Like thesensing layer 9 the enzyme-free layer 10 was applied as a paste. Thatpaste differs from the paste of the sensing layer 9 only in that noenzyme was added to it.

The diffusion barrier also comprises a layer 11 which prevents largemolecules from clogging pores of the sensing layer 9. The layer 11 maybe a dialysis layer which can be provided as a membrane made ofcellulose and/or a polymer material. Such a dialysis layer is also anenzyme-free layer and may be applied directly on top of the sensinglayer 9 or, as shown in FIG. 1, on top of the electrically conductiveenzyme-free layer 10. In one embodiment, such a dialysis layer shouldhinder analyte diffusion as little as possible. In other embodiments,the layer 11 has an effective diffusion coefficient D_(eff) for theanalyte which is at most about ten times lower than the diffusioncoefficient D of the analyte in water, and in yet other embodiments itis at most about five times lower than the diffusion coefficient D ofthe analyte in water. A dialysis layer can be applied as a solid film orapplied as a polymer solution which hardens into a dialysis membranein-situ.

Dialysis membranes are often characterized by their molecular weight cutoff (MWCO) which depends on the pore size. The MWCO describes themolecular weight at which a compound will be 90% retained following anight (17-hour) dialysis. The dialysis layer of the example shown has aMWCO of less than about 10 kDalton, and may be less than about 7 kD, andeven less than about 5 kD in some embodiments. It has to be understoodthat MWCOs stated for dialysis layers apply strictly only to globularmolecules such as most proteins. More linear molecules may be able topass through the pores of a dialysis layer, even if their molecularweight exceeds the stated MWCO.

Instead of or in addition to a dialysis membrane, the diffusion barriermay also comprise a polymer layer made of a polymer having azwitterionic structure to protect the sensing layer 9 and any porouslayer 10 from ingression of proteins. A zwitterionic structure enablesthe rapid uptake of polar protic solvents, in particular water, and suchanalytes as glucose dissolved within. Hence, polymers having azwitterionic structure attached to a polymeric backbone are impermeablefor proteins but hinder diffusion of analytes like glucose very little.A well-known example for such a polymer is poly(2-methacryloyloxyethylphosphorylcholine-co-n-butylmethacrylat) (MPC for short). The MPCpolymer layer 11 is applied as a polymer solution comprising ethanol ordistilled water and at least 5 wt. % MPC, especially at least 10 wt. %MPC.

In these exemplary embodiments, the diffusion barrier and especially thepolymer layer 11, which it comprises, protect the sensor 1 frommechanical damage during the implantation process, prevent enzyme fromleaking out of the sensing layer 9 into surrounding tissue, where itmight be harmful, and prevents large molecules from clogging pores ofthe sensing layer 9. It is possible to mix a polymer having azwitterionic structure like MPC with another polymer, for examplepolyurethane or typical constituents of the above-mentioned dialysemembranes, in order to tune physical properties of the polymer layer 11.

It is also possible to tune physical properties, such as thepermeability for analytes, of the layer 11, if this contains aco-polymer having constituents of different hydrophilicity, by varyingthe relative amount of each constituent in the co-polymer. In the caseof MPC it is possible to increase the relative amount of2-methacryloyloxyethyl phosphorylcholine versus that of butylmethacrylatfrom 30:70% to 50:50% yielding a co-polymer with higher permeabilitytowards polar protic solvents or glucose. Another way to increasepermeability for polar protic solvent or glucose is to change thehydrophobic backbone of the co-polymer into a more hydrophilic entity.This also applies for other water-soluble analytes.

The sensing layer 9 of the example shown in FIG. 1 contains porousparticles 13 to increase its porosity and thereby ease diffusion ofanalyte molecules into the sensing layer 9. Porous particles 13 in thisrespect are particles which have voids to adsorb water molecules. Theseporous particles 13 are added to the paste from which the sensing layer9 is formed and cause voids through which analyte molecules and watermay pass. The porous particles 13 are bound with other particles of thepaste by the polymeric binder. Carbon nanotubes are an especially usefuladditive to increase the porosity of the sensing layer as they tend toform clews, which are only partially filled with carbon particles andbinder, and also increase the electrical conductivity of the sensinglayer. Silica particles may also be used as porous particles 13 toincrease the porosity of the sensing layer 9.

If silica or similar porous particles 13 are used, it is advantageous touse material with a particle size distribution such that the maximumparticle size is less than the thickness of the sensing layer 9. In oneembodiment, the porous particles 13 should measure at least about 1 μm,and in other embodiments should measure at least about 5 μm. Consideringa sensing layer 9 thickness around about 20 μm to about 50 μm, silica FK320 from Degussa provides adequate particle size, up to about 15 μm.Typically, less than about 10% of this material are mixed into thepaste, and can be as low as about 5% or less.

In one embodiment, electrical conductivity is provided throughout thesensing layer 9, so that at each spot of the porous matrix, where aproduct molecule is generated from the enzymatic reaction, this moleculeis directly oxidized or reduced by application of the appropriatevoltage without the need for extended diffusion of this molecule to adistant site. Under these circumstances the porous and permeable sensinglayer 9 is capable of electrolyzing the analyte substantially throughoutthe entire layer.

Whatever means for increasing the porosity is used, the mixing of theenzyme with the paste will lead to a fraction of enzyme molecules beingaccessible to the analyte, either on the upper surface of the sensinglayer 9, or at the channels in the vicinity of the additive particleswithin the sensing layer 9. In one embodiment, the enzyme is immobilizedby adsorption and entrapment in the working electrode 3. Entrapmentdepends not only on the sensing layer 9 but also on properties of thediffusion barrier, i.e. the layer 11, and of the optional enzyme-freelayer 10. It is understood that in order to maintain the idealdistribution of enzyme within the working electrode, contact withsolvent (water) should not lead to massive detachment of enzyme from thematrix and subsequent migration of enzyme molecules. In otherembodiments, enzyme immobilization in the sensing layer 9 can beenhanced by cross-linking. Especially advantageous are enzyme moleculeswhich are cross-linked as a chain, keeping in mind, however, that longerchains result in the enzyme being less effective. In one embodiment,therefore, on average about three to about ten enzyme molecules arelinked together, and in other embodiments about 4 to about 8 on averageare linked, wherein an average chain length of about five to about sevenenzyme molecules is typically useful.

It is possible to add a cross-linking agent, i.e. glutaraldehydsolution, to the paste before drying. However, it is also possible tomix an already cross-linked enzyme into the paste. In one embodiment, anenzyme is used which forms a complex with a hydrophilic partner. Afterbeing mixed into a paste which is less hydrophilic or even hydrophobic,as can be achieved by mixing carbon particles with suitable binders, thecross-linked enzyme sits in a local hydrophilic environment whichcontributes to its stability. One result of a cross-linked enzyme with ahydrophilic partner is that it enhances migration of hydrated analytemolecules towards the enzyme. Thus the wetting of the sensing layer 9 isaccelerated which shortens the wet-up time of the sensor afterimplantation. As a specific example, glucose oxidase cross-linked withdextrane from Roche Diagnostics (Penzberg, Germany, Ident-No.1485938001) has been found to have such a content of enzyme(approximately 16%) that enough activity (20 to 40 U/mg lyophylisate)can be preserved.

By mixing already cross-linked enzyme with the sensing layer pastecontaining carbon nanotubes, the trait of the carbon nanotubes to windup and form clews, which act as macroporous cage structures, issupported by the larger enzyme-dextrane chains, in particular by theiraggregation. As a consequence, the cross-linking enzyme will assist inthe formation of porous structures of the sensing layer 9.

The sensing layer 9 of the example shown comprises carbon particles withan average size of less than about 1 μm, polymeric binder, enzyme andcarbon nanotubes as porous particles 13. The porous particles 13 aremore effective to increase the porosity of the sensing layer 9 if theyare significantly larger than the carbon particles. In the exampleshown, the porous particles 13 measure at least about 1 μm, and can beat least about 5 μm, on average. Typically the sensing layer 9 comprisesabout 50 wt. % to about 70 wt. % polymeric binder, about 20 wt % toabout 40 wt. % carbon particles and up to about 20 wt. % (and as low asbetween about 1 wt. % and about 10 wt. %) porous particles 13 likecarbon nanotubes or silica. Carbon nanotubes are an especiallyadvantageous additive as they increase both the porosity and theelectrical conductivity of the sensing layer 9. In the embodiment shownschematically in FIG. 1, multiwall carbon nanotubes (research grade,purity>95%) by NanoLab, Newton, Mass., of length about 5 μm to about 20μm and an average outer diameter of about 25 nm to about 35 nm have beenused. The binder is a thermoplastic resin, e.g. on the basis of an epoxyresin or on the basis of polyvinyl chloride (PVC)/polyvinyl alcohol(PVA). Resins on the basis of a fluoro carbon resin, particularlypolytetrafluoroethylene, or of polystyrene, may also be used as binders.In the case of PVC/PVA binders, the use of additives like silicone oilcan help to adjust the viscosity of the paste.

In this way the sensing layer 9 of the sensor 1 shown in FIG. 1 isadapted and arranged in such a way that in operation after implantationthe analyte concentration in the sensing layer 9 is highest at the uppersurface, decreases with increasing distance from the upper surface, andis zero at the lower surface which is the furthest point from theanalyte-containing body fluid and which touches the contact pad 7. Theenzyme loading of the sensing layer 9, i.e. the amount of the enzymeimmobilized therein, should be chosen with respect to the porosity andwater-permeability of the sensing layer 9.

An important parameter in this respect is the effective diffusioncoefficient D_(eff) of the sensing layer 9. The effective diffusioncoefficient D_(eff) characterizes the diffusion of the analyte in thesensing layer 9 and depends on the pore volume ε and the tortuosity τ ofthe sensing layer 9. Generally, the effective diffusion coefficientD_(eff) can be described as D_(eff)=D·ε/τ, wherein D is the diffusioncoefficient of the analyte in water. The quotient τ/ε is also calledhindrance H. In the example shown H is between about 10 and about 1000,and can be between about 50 and about 500.

Another important parameter in this respect is the enzyme loadingparameter α which can be described as α=(V_(max)·d)/(K_(M)·D) whereinV_(max) is the enzyme activity density which determines the maximumspeed of analyte conversion, K_(M) the Michaelis Menten constant of theenzyme, d the thickness of the sensing layer and D the diffusioncoefficient of the analyte in water. In one embodiment, the ratio of theeffective diffusion coefficient D_(eff) in the sensing layer 9 and theenzyme loading parameter α is in the range of about 10 to about 200.

FIG. 2 shows the functional characteristic of the sensor describedabove. The layer 11 has been made from MPC (Lipidure CM 5206, NOF Corp.Japan) by dispensing a 10% solution of MPC in ethanol/water on theelectrodes. The measurement current I in nA is plotted versus glucoseconcentration g in mg/dl. The data shown in FIG. 2 were measuredin-vitro in aqueous glucose solution. As can be seen, no saturation athigher glucose concentrations is observed.

FIG. 3 shows for comparison measurement currents I_(A) and I_(B) in nA.Currents I_(A) were measured in-vitro, currents I_(B) by a sensor in abiomatrix, both at a temperature T=35° C., after the sensors had beenequilibrating in the respective medium for 12 hours. Every data pointshown belongs thus to a biomatrix measurement and an aqueous glucosesolution measurement at identical glucose concentration. The biomatrixused consists of stabilized blood plasma, to which glucose was added inorder to obtain the desired glucose concentration. Sensor currentsmeasured in the biomatrix and sensor currents measured in aqueousglucose solution show excellent agreement.

The result is particular noteworthy and demonstrates the profoundeffects arising from the sensor layout of the embodiment of the presentinvention. In general, it is expected, that exposure of a sensor 1 to abiomatrix ensures the deposition of proteins, peptides or fibrin on thesensor surface. This process affects the permeability for analytes orwater of an outer layer, such as layer 11. In a conventional sensorlayout, this layer restricts diffusion of analyte to the sensing layerso that a permeability decrease results in a weaker measurement signal.

However, the signal height of the described sensor 1 is not affected byexposure to a biomatrix, as seen in FIG. 3, since the diffusion ofanalyte through the layer 11 is not the rate-limiting step in thegeneration of the signal. Therefore any permeability alteration has verylittle effect on the signal for the sensor 1 described above.

The present invention is not limited to enzymes using oxygen asco-substrate in the catalytic reaction. The enzyme may be adehydrogenase as well. For example, a glucose dehydrogenase which doesnot use oxygen as co-substrate can be distributed within the sensinglayer 9. Known dehydrogenases include certain molecules as cofactors forthe oxidation of glucose, for example pyrroloquinoline quinone (PQQ), orflavin adenine dinucleotide (FAD) or nicotinamid adenine dinucleotide(NAD), see EP 1 661 516 A1. Any of these dehydrogenases can be used inthe sensing layer 9 instead of an oxidase.

FIG. 4 shows the sensor current I in nA, which was measured by sensors Ato J in phosphate-buffered aqueous glucose solution of differentconcentrations. Sensor currents measured at a glucose concentration of360 mg/dl are depicted by triangles (▴). Sensor currents measured at aglucose concentration of 180 mg/dl are depicted by squares (▪). Sensorcurrents at a glucose concentration of zero are depicted by diamonds(♦).

Sensors A to J differ only with respect to the diffusion barrier appliedon top of the sensing layer 9. At sensors A to E a diffusion barrier islacking, i.e. the sensing layer 9 is in direct contact with the aqueousglucose solution to be measured. Sensors F to J comprise a diffusionbarrier covering the sensing layer 9. The diffusion barrier of sensors Fto J was provided as a polymer layer made of MPC-like layer 11 inFIG. 1. As can be seen, the sensor currents are only slightly lower forsensors F to J than for sensors A to E. Hence, the diffusion barrierprovided by the MPC polymer layer 11 hinders diffusion of analytemolecules only to a very small extent. As the sensor currents of sensorsF to J are about 20% lower than sensor currents of sensors A to E, itcan be concluded that the diffusion barrier of sensors F to J leads toanalyte concentrations on the upper surface of the sensing layer 9 whichare only about 20% lower than in the glucose solution surrounding thesensors.

As in the embodiment described previously with reference to FIG. 1 thesensing layers 9 of sensors A to J comprise cross-linked enzymes, i.e.dextranized glucose oxidase which can be purchased by Roche Diagnostics,Penzberg, Germany, Ident-No. 14859389001. The dextranized glucoseoxidase has been dissolved in phosphate-buffered solution and mixed intoa paste comprising carbon particles, carbon nanotubes and polymericbinder. The sensing layer 9 was dispensed on the contact pad 7 of theworking electrode 3 on the sensor substrate 5 with a spot size of about0.05 mm² to about 0.1 mm², e.g. a circular spot of about 300 μmdiameter. The thickness of the sensing layer 9 was about 20 μm. AnAg/AgCl reference electrode 4 of consistent size has also been provided.The counter electrode 2 was of rectangular shape (approximately 400 μmby 900 μm) with an about 20 μm thick layer of carbon paste containingcarbon nanotubes.

It is seen in FIG. 4 that the sensor current is barely affected by thepresence of the membrane made from MPC.

It can be concluded from this finding that, by the particular choice ofthe zwitterionic membrane structure, a coating which is highly permeablefor solvated glucose has been found. For the construction of a sensor 1,where diffusion-limitation occurs in the sensing layer 9 (see FIG. 1),such a high permeability of the membrane is important. Conversely, thediffusion of analyte through the diffusion barrier provided by theMPC-layer 11, should be hindered as little as possible, ideally theanalyte concentration (i.e. signal) at the sensing layer 9 with thecoating should be not less than half of the value obtained without thecoating.

It should be noted that the optional enzyme-free layer 10 should alsohave little hindrance to analyte diffusion, therefore its layerthickness should be much thinner than that of the sensing layer 9.

As stated before, mixing of a hydrophilized cross-linked enzyme canyield a very stable function over extended periods of time, since wet-upof the sensing layer is fast and the enzyme distribution stays constant.This is reflected by drift values obtained in measuring the abovesensors over 6 days in aqueous glucose solution. For the uncoatedsensors, drift ranges from about −0.62% per day to about 0.78% per day,while the coated ones cover a range from about −0.5% to about 1.5% perday. These small drift values have been measured at 37° C.

The particular advantage of measurement stability, i.e. low signaldrift, is not limited to a sensor 1 with an enzyme in the sensing layer9 which uses oxygen as a co-substrate. In fact, the same benefit ofcross-linking can be obtained by using a cross-linked dehydrogenasewhich does not need oxygen as a co-substrate in the catalytic reaction.For example, a dextranized glucose dehydrogenase or a pegylateddehydrogenase (PEG: polyethylene glycol) can be brought into the sensinglayer 9.

In the sensor 1 shown in FIG. 1 the sensing layer 9 is arranged on thecontact pad 7. Further, the sensing layer 9 has a lower surface facingthe contact pad 7 and an upper surface facing away from the contact pad7, or more generally the sensing layer 9 has having a lower surfacefacing the support member 5 and an upper surface facing away from thesupport member 5 toward the analyte-containing body fluid.Correspondingly the layers 12, 14 are arranged the contact pads 6, 8.FIG. 5 shows an amended embodiment of the sensor 1 of FIG. 1. Theembodiment of FIG. 5 corresponds to the embodiment of FIG. 1, with thedifference that the electrodes with the contact pads 6, 7, 8 beingplaced on the side of the water-permeable layers 9, 12 and 14, incontrast to FIG. 1. It is also possible to place a contact pad 6, 7, 8on two sides of the respective layer 9, 12, 14, as illustrated for thecontact pad 6 of the water-permeable layer 12 of counter electrode 2.This contact pad 6 can also be formed so that it encloses the layer 12from all sides. In all cases where the contact pad 6, 7, 8 sits on theside of the permeable layer 9, 12, 14, the surface of the layer 9, 12,14 facing away from the analyte-containing body fluid is directly incontact with the support member 5.

The features disclosed in the above description, the claims and thedrawings may be important both individually and in any combination withone another for implementing the invention in its various embodiments.

It is noted that terms like “preferably”, “commonly”, and “typically”are not utilized herein to limit the scope of the claimed invention orto imply that certain features are critical, essential, or even,important to the structure or function of the claimed invention. Rather,these terms are merely intended to highlight alternative or additionalfeatures that may or may not be utilized in a particular embodiment ofthe present invention.

For the purposes of describing and defining the present invention it isnoted that the term “substantially” is utilized herein to represent theinherent degree of uncertainty that may be attributed to anyquantitative comparison, value, measurement, or other representation.The term “substantially” is also utilized herein to represent the degreeby which a quantitative representation may vary from a stated referencewithout resulting in a change in the basic function of the subjectmatter at issue.

Having described the present invention in detail and by reference tospecific embodiments thereof, it will be apparent that modification andvariations are possible without departing from the scope of the presentinvention defined in the appended claims. More specifically, althoughsome aspects of the present invention are identified herein as preferredor particularly advantageous, it is contemplated that the presentinvention is not necessarily limited to these preferred aspects of thepresent invention.

1. An amperometric sensor configured for implantation into the livingbody of a human or animal to measure the concentration of an analyte ina body fluid, said sensor comprising a counter electrode and a workingelectrode, said working electrode comprising a sensing layer which isgenerally permeable for water and arranged on a support member adjacentto a contact pad, said sensing layer comprising an immobilized enzymecapable of acting catalytically in the presence of the analyte to causean electrical signal, the sensing layer having an upper surface facingthe body fluid and a lower surface facing away from the body fluid,wherein the immobilized enzyme is distributed in the sensing layer insuch a way that the enzyme concentration in the middle between the upperand lower surfaces is at least as high as on the upper surface of thesensing layer.
 2. The sensor according to claim 1, wherein the sensinglayer has an effective diffusion coefficient D_(eff), whichcharacterizes the diffusion of the analyte in the sensing layer and isabout 10-times to about 1000-times lower than the diffusion coefficientD of the analyte in water.
 3. The sensor according to claim 1, whereinthe enzyme comprises a cross-linked enzyme.
 4. The sensor according toclaim 3, wherein the cross-linked enzyme forms a complex with ahydrophilic partner, and wherein the cross-linked enzyme is an enzyme inwhich the enzyme molecules form a cross-linked complex with ahydrophilic partner.
 5. The sensor according to claim 3, wherein thecross-linked enzyme has an average chain length of about three to aboutenzyme molecules.
 6. The sensor according to claim 5, wherein theaverage chain length is about 4 to about 8 enzyme molecules.
 7. Thesensor according to claim 1, wherein the sensing layer is covered by adiffusion barrier which hinders diffusion of analyte molecules to suchan extent that after implantation into the living body of a human or ananimal the analyte concentration at the upper surface of the sensinglayer is at most about ten times lower than in the body fluidsurrounding the implanted sensor and wherein the diffusion barriercomprises an electrically conductive enzyme-free layer comprising carbonparticles and a polymeric binder.
 8. The sensor according to claim 7,wherein after implantation into the living body of a human or an animalthe analyte concentration at the upper surface of the sensing layer isat most about five times lower than in the body fluid surrounding theimplanted sensor.
 9. The sensor according to claim 1, wherein thesensing layer contains porous particles.
 10. The sensor according toclaim 9, wherein the porous particles comprise one or both of silica andcarbon nanotubes.
 11. The sensor according to claim 9, wherein theporous particles increase both the porosity and the electricalconductivity of the sensing layer.
 12. The sensor according to claim 9,wherein the porous particles measure on average between at least about 1μm and about 5 μm.
 13. The sensor according to claim 1, wherein thesensing layer is adapted and arranged in such a way that in operationafter implantation the analyte concentration in the sensing layer ishighest at the upper surface, decreases with increasing distance fromthe upper surface, and is zero at the lower surface which is thefurthest point from the analyte-containing body fluid.
 14. The sensoraccording to claim 1, wherein the contact pad of the working electrodeis an electrically conductive film.
 15. The sensor according to claim14, wherein the electrically conductive film comprises a metallic filmor a polymer film.
 16. The sensor according to claim 1, wherein theworking electrode is arranged on a support member made of a plasticmaterial.
 17. The sensor according to claim 1, wherein the sensing layerhas a thickness of between at least about 5 μm and about 100 μm.
 18. Thesensor according to claim 1, wherein the sensing layer has a thicknessof between at least about 20 μm and about 70 μm.
 19. The sensoraccording to claim 1, wherein the sensing layer comprises carbonparticles and a polymeric binder.
 20. The sensor according to claim 1,wherein the sensing layer is covered by a diffusion barrier whichhinders diffusion of analyte molecules to such an extent that afterimplantation into the living body of a human or an animal the analyteconcentration at the upper surface of the sensing layer is at most aboutten times lower than in the body fluid surrounding the implanted sensor.21. The sensor according to claim 20, wherein the sensing layer iscovered by a diffusion barrier which hinders diffusion of analytemolecules to such an extent that after implantation into the living bodyof a human or an animal the analyte concentration at the upper surfaceof the sensing layer is at most about five times lower than in the bodyfluid surrounding the implanted sensor.
 22. The sensor according toclaim 20, wherein the diffusion barrier comprises a dialysis layer. 23.The sensor according to claim 20, wherein the diffusion barriercomprises a polymer layer made of a polymer having a zwitterionicstructure.
 24. The sensor according to claim 1, wherein the sensinglayer includes lateral surfaces which are generally impervious for thebody fluid.
 25. The sensor according to claim 1, wherein the sensinglayer is electrically conductive.
 26. The sensor according to claim 25,wherein the sensing layer has an electrical conductivity of at leastabout 1 Ω⁻¹ cm⁻¹.
 27. The sensor according to claim 1, wherein theenzyme comprises an oxidase.
 28. The sensor according to claim 1,wherein the enzyme comprises a dehydrogenase.
 29. The sensor accordingto claim 1, wherein the enzyme is distributed generally equallythroughout the sensing layer.
 30. The sensor according to claim 29,wherein the sensing layer has an enzyme loading parameter α which isbetween about 10 to about 200 times smaller than the effective diffusioncoefficient D_(eff).
 31. The sensor according to claim 1, wherein saidsensing layer is arranged on the contact pad and the sensing layer has alower surface facing the contact pad and an upper surface facing awayfrom the contact pad.
 32. A method for manufacturing an amperometricsensor configured for implantation into the living body of a human oranimal to measure the concentration of an analyte in a body fluid, saidmethod comprising the following steps: mixing carbon particles, enzymeand a polymeric binder to create a paste; applying said paste adjacentto a contact pad on a support member of said sensor, hardening thatpaste into a porous sensing layer.